In endoscopic imaging systems, high image resolution and high sensitivity (or low radiation) is an important system characteristic. This is particularly true in medical imaging through where the clarity and contrasts within an image directly affect the diagnostic capabilities of a physician. That is, the higher the resolution and the higher the sensitivity, the earlier and easier the detection of abnormalities is. Likewise, industrial uses such as quality control of product components operate in much the same manner and lack of detection of abnormalities can have similarly disastrous results.
Higher resolution images at low radiation levels can also help distinguish aspects of the image thus presenting additional valuable information. For example, if the image shows certain fractal duct structures then a physician may be able to deduce that a tumor is benign. Further, accurate representation of objects in the image, as to image size, for example, assists in diagnosis. That is, the observation of a stable tumor size over time alleviates the fear of malignancies without intrusive and invasive operations.
Applications of fluorescent endoscopy compare tissue regions based upon different image signals for at least two light wavelengths, e.g. red and green images for the same tissue. However, the process is typically limited by detector noise and the rate at which the different color images are provided. The noise is typically at the level of 100 electrons, while the rate of acquiring a red and green image pair is as much as a few hundred milliseconds or more, requiring patient immobilization and mechanically fixed, i.e. tripod-mounted endoscope probes which seriously inhibit the usability of present day fluorescent endoscopes. The latter could be solved by using a pair of photo detectors assemblies for measuring the intensity of light emitted from the tissue at two different wavelengths simultaneously. However, two detectors are clumsy and requires expensive cross-correlation which is costly in time and expense.
A proximity electron bombardment charge-coupled device (EBCCD) has no focussing. To have acceptable resolution, the distance between the EBCCD chip and photocathode must be kept small (0.7–0.8 mm). At this distance, it is impossible to apply high voltage to the elements of the EBCCD. The maximum gain is typically limited to a few hundred. Moreover, positive ions are produced by energetic electrons which after hitting the anode, strike back at the cathode to produce spurious electrons, which hit the anode again at different positions from the original electron. Such spurious electrons produce noise in the image signal, destroys the image resolution and shorten the life of the EBCCD. Therefore, intensified image CCDs have heretofore included micro-channel-plate (MCP) elements to reduce these noise source. However, the MCP itself is noisy with only about 60% acceptance.
Prior intensified CCD devices all use the (light-spreading) phosphor screen as the transmitting medium and the (noisy and with small acceptance) MCP as the amplification means. All such intensified devices have poor modulate transfer functions (MTF) and poor image quality.
Medical diagnostic imaging systems utilizing x-ray image intensifier tubes are well known in the art. The image intensifier tube has as a component a scintillator that converts an x-ray image, representing the absorption of x-rays by the structure to be depicted, into visible light. Devices such as this are widely used for medical observation. The visible light can then be made to impinge upon a photographic film or a photosensitive detector that electronically records the image. The film can then be developed for direct review, at the expense of time, or the electronic signals from the detector can be processed and transmitted to a cathode-ray tube (‘CRT’) or photographic recording system.
FIG. 1 shows a prior art scintillator 10, which is generally formed by depositing cesium iodide by vacuum evaporation onto a substrate 14. The thickness of the cesium iodide, or structured cesium, deposited generally ranges from 150–500 microns. The cesium iodide is deposited in the form of needles 12 each with a diameter of 5–10 microns. Since the refractive index of cesium iodide is 1.8, a fiber optic effect is obtained. This effect minimizes the lateral diffusion of the light within the scintillating material. A scintillator of this type, for example, is described in U.S. Pat. No. 4,803,366 dated Feb. 7, 1989.
The resolution of the image intensifier tube depends on the capacity of the cesium iodide needles 12 to properly channel the light. Non-uniformity (i.e. positions dependent light yield across the needles) and cross talks between the needles can result in large non-Gaussian tails which degrade the spatial image resolution. The cesium iodide as well as another popular material, sodium iodide, used as x-ray converters all have relatively low densities and thus low detective quantum efficiency (“DQE) if a thin layer of scintillator is used, and/or poor spatial resolution if a thick layer of scintillator is used.
These factors can be seen with more particularity in FIG. 2 which shows the blooming of a single pixel imaged using these conventional scintillators. The vertical axis represents intensity of the pixel and the horizontal axis represents position relative to the center of the pixel with respect to light. One skilled in the art will understand that the broader a particular function of light for a pixel appears on this graph, the lower potential resolution on a photosensitive medium, such as film or a CRT, since this will represent a blooming and a potential for cross-talk between individual pixels. Each line represents different prior art systems. Line 20 represents a Lanex fast screen; line 22 represents a non-structured cesium iodide crystal layer of 220 Micron thickness; line 24 represents a structured cesium iodide layer of 220 micron thickness; line 26 represents a Lanex fine screen; and line 28 represents a structured cesium iodide layer of 75 micron thickness.
Often, as is the case with x-rays, the radiation used to create the image has potentially harmful effects on the subject of the examination. Devices with higher DQE reduce the required radiation doses per viewing and allow more frequent viewings for the observation of the growth rate of abnormalities. The density of cesium iodide and sodium iodide crystals is low, thus, prior art scintillators have a low DQE when the scintillator is thin. DQE can be raised by increasing the thickness, but this is done at the expense of spatial resolution.
Conventional methods of fabrication of scintillators, such as vacuum deposition or chemical vapor deposition, have difficulty making films of single crystals of more than a few microns thick. This, in turn, detrimentally affects the light conversion efficiency of the scintillator.
Once the scintillator converts the x-ray image into visible light, there is often still the problem of inadequate light to adequately resolve objects clearly in the image by a detector in the image intensifier tube. The problem is common in various other applications such as endoscopic or laparoscopic imaging, and non-medical imaging such as night-vision photography, for example. Commercially available systems of the aforementioned types generally use as a detector a room temperature charge-coupled device (“CCD”) to electronically capture the image-bearing light. Such a CCD has no gain and, therefore, low signal-to-noise ratio, thus requiring intense light illumination. Each pixel in the CCD converts incoming photons into electron-hole pairs. This conversion is made with an efficiency about 30%. Mainly due to the thermal noise of the readout electronics, there is a large noise produced in each pixel even if there is no input light. This noise is typically 100 electrons per pixel for 10 MHz readout frequency. Therefore, in order to have a reasonable signal-to-noise ratio, about 2000 photons per pixel are needed for a standard CCD at room temperature, with a quantum efficiency 30%.
One solution to this problem has been to use a cooled CCD, which has less noise because it is cooled to a low temperature. Even with the cooled CCD though, a large quantity of photons, approximately 400 photons, per pixel are required to have a reasonable signal-to-noise ratio. Cooled CCDs usually have slow read-out speed.
Previously implemented proximity electron-bombarded CCDs achieve some of the desired sensitivity but have other drawbacks, such as short lifetimes, low gain and high noise levels. Moreover, their image size is limited to be the same as the size of the CCD, thus very small for any practical uses like mammogram.
Implementation with Complementary Metal-Oxide Semiconductors (CMOS) sensors would be less promising, as CMOS sensors have much bigger “noise,” the industry's term for annoying little dots or scratches on photos, than even usual CCDS, let alone EBCCDS, and thus even worse pictures in sharpness, quality and sensitivity.
In medical imaging, a further diagnostic advantage is gained by three-dimensional reconstruction of images. Such reconstruction followed by reproduction on the screen of the various representations of tissue, such as the breast, density (similar to representation available in computer tomography) has great clinical values, facilitating the diagnostics and reducing the percentage of errors. However, prior art systems have not been able to reproduce exact spatial fixation of the soft flexible tissue, which is too flexible for its fixation with submillimeter accuracy. Therefore, image shadows do not match accurately enough to allow the reconstruction of the three-dimensional image with the full resolution of the detector.
Accordingly, it is an object of the invention to provide a fluorescence endoscope with reduced noise and improved sensitivity, resolution and speed of image acquisition to allow the probe and the patient to move nominally without smearing out the images taken at different wavelengths with a single photo-detector.
It is another object of this invention to provide a scintillator that resists blooming and pixel cross talk so as to create a high-resolution image at lower radiation doses.
It is still another object of this invention to provide a scintillator with a high DQE without sacrificing spatial resolution.
It is still another object of this invention to provide a scintillator of high resolution that can be used for both displaying an image on an electronic screen and presenting the image to photographic film.
It is still another object of the invention to provide a CCD and a CMOS detector that operates at low levels of light.
It is still another object of the invention to provide an accurate three-dimensional image.
It is a further object of the invention to provide methods in accord with the above apparatus.
These and other objects of the invention will appear with the descriptions and exemplary illustrations provided hereinafter.